Split energy level radiation detection

ABSTRACT

An energy discriminating apparatus and method is disclosed for use in connection with digital radiography and fluoroscopy. In use of the detection system and method an x-ray source is actuated to direct x-rays through a patient&#39;s body, the x-rays including both higher and lower energy radiation. A first detector element, including a plurality of segments, is positioned opposite the source to receive and respond predominantly to x-rays in a lower energy range, the remaining x-rays, being generally of higher energy, passing through the first detector element. A second detector element, also including a plurality of segments, each segment including a phosphor coating layer and a sensor, is positioned to receive and respond to the higher energy radiation passing through the first element. The sensors are coupled respectively to each detector element segment for substantially simultaneously sensing the response and spatial location, relative to the detector elements, of radiation to which each detector element respectively responds. A filter element is interposed between the first and second detectors to enhance discrimination in the energy response of the respective detector elements. Particular preferred detector phosphor materials are identified. The sensors produce separately and simultaneously information representing patterns of relatively lower and higher energy emergent from the patient&#39;s body. Digital data processing and conversion equipment responds to the sensors to produce digital information representing each of said images, which can be digitally processed to enhance image characteristics.

DESCRIPTION

1. Technical Field

This invention relates to the field of medical diagnostic imaging andmore particularly to an improved x-ray detector for use in digitalradiography and fluoroscopy. The detector provides separate simultaneousrepresentations of different energy radiation emergent from a subject.

2. Background Art

Radiography and fluoroscopy are long well known diagnostic imagingtechniques.

In a conventional radiography system, an x-ray source is actuated todirect a divergent area beam of x-rays through a patient. A cassettecontaining an x-ray sensitive phosphor screen and film is positioned inthe x-ray path on the side of the patient opposite the source. Radiationpassing through the patient's body is attenuated in varying degrees inaccordance with the various types of tissue through which the x-rayspass. The attenuated x-rays from the patient emerge in a pattern, andstrike the phosphor screen, which in turn exposes the film. The x-rayfilm is processed to yield a visible image which can be interpreted by aradiologist as defining internal body structure and/or condition of thepatient.

In conventional fluoroscopy, a continuous or rapidly pulsed area beam ofx-rays is directed through the patient's body. An image intensifier tubeis positioned in the path of the beam opposite the source with respectto the patient. The image intensifier tube receives the emergentradiation pattern from the patient, and converts it to a small,brightened visible image at an output face. Either a mirror or closedcircuit television system views the output face and produces a dynamicreal time visual image, such as on a CRT, a visual image forinterpretation by a radiologist.

More recently, digital radiography and fluoroscopy techniques have beendeveloped. In digital radiography, the source directs x-radiationthrough a patient's body to a detector in the beam path beyond thepatient. The detector, by use of appropriate sensor means, responds toincident radiation to produce analog signals representing the sensedradiation image, which signals are converted to digital information andfed to a digital data processing unit. The data processing unit records,and/or processes and enhances the digital data. A display unit respondsto the appropriate digital data representing the image to convert thedigital information back into analog form and produce a visual displayof the patient's internal body structure derived from the acquired imagepattern of radiation emergent from the patient's body. The displaysystem can be coupled directly to the digital data processing unit forsubstantially real time imaging, or can be fed stored digital data fromdigital storage means such as tapes or discs representing patient imagesfrom earlier studies.

Digital radiography includes radiographic techniques in which a thin fanbeam of x-ray is used, and other techniques in which a more widelydispersed so-called “area beam” is used. In the former technique, oftencalled “scan (or slit) projection radiography” (SPR) a fan beam of x-rayis directed through a patient's body. The fan is scanned across to thepatient, or the patient is movably interposed between the fan beam x-raysource and an array of individual cellular detector segments which arealigned along an arcuate or linear path. Relative movement is effectedbetween the source-detector arrangement and the patient's body, keepingthe detector aligned with the beam, such that a large area of thepatient's body is scanned by the fan beam of x-rays. Each of thedetector segments produces analog signals indicating characteristics ofthe received x-rays.

These analog signals are digitized and fed to a data processing unitwhich operates on the data in a predetermined fashion to actuate displayapparatus to produce a display image representing the internal structureand/or condition of the patient's body.

In use of the “area” beam, a divergent beam of x-ray is directed throughthe patient's body toward the input face of an image intensifier tubepositioned opposite the patient with respect to the source. The tubeoutput face is viewed by a television camera. The camera video signal isdigitized, fed to a data processing unit, and subsequently converted toa tangible representation of the patient's internal body structure orcondition.

One of the advantages of digital radiography and fluoroscopy is that thedigital image information generated from the emergent radiation patternincident on the detector can be processed, more easily than analog data,in various ways to enhance certain aspects of the image, to make theimage more readily intelligible and to display a wider range ofanatomical attenuation differences.

An important technique for enhancing a digitally represented image iscalled “subtraction”. There are two types of subtraction techniques, onebeing “temporal” substraction, the other “energy” subtraction.

Temporal, sometimes called “mask mode” subtraction, is a technique thatcan be used to remove overlying and underlying structures from an imagewhen the object of interest is enhanced by a radiopaque contrast agent,administered intra-arterially or intra-venously. Images are acquiredwith and without the contrast agent present and the data representingthe former image is subtracted from the data representing the latter,substantially cancelling out all but the blood vessels or anatomicalregions containing the contrast agent. Temporal subtraction is,theoretically, the optimum way to image the enhancement caused by anadministered contrast agent. It “pulls” the affected regions out of aninterfering background.

A principle limitation of digital temporal subtraction is thesusceptibility to misregistration, or “motion” artifacts caused bypatient movement between the acquisition of the images with and withoutthe contrast agent.

Another disadvantage of temporal subtraction is that it requires the useof a contrast material and changes in the contrast caused by the agentmust occur rapidly, to minimize the occurrence of motion causedartifacts by reducing the time between the first and second exposureacquisition. Temporal subtraction is also not useful in studiesinvolving rapidly moving organs such as the heart. Also, theadministration of contrast agents is contraindicated in some patients.

An alternative to temporal subtraction, which is less susceptible tomotion artifacts, is energy subtraction Whereas temporal subtractiondepends on changes in the contrast distribution with time, energysubtraction exploits energy-related differences in attenuationproperties of various types of tissue, such as soft tissue and bone.

It is known that different tissues, such as soft tissue (which is mostlywater) and bone, exhibit different characteristics in their capabilitiesto attenuate x-radiation of differing energy levels.

It is also known that the capability of soft tissue to attenuatex-radiation is less dependent on the x-ray's energy level than is thecapability of bone to attenuate x-rays. Soft tissue shows less change inattenuation capability with respect to energy than does bone.

This phenomenon enables performance of energy subtraction. In practicingthat technique, pulses of x-rays having alternating higher and lowerenergy levels are directed through the patient's body. When a lowerenergy pulse is so generated, the detector and associated digitalprocessing unit cooperate to acquire and store a set of digital datarepresenting the image produced in response to the lower energy pulse. Avery short time later, when the higher energy pulse is produced, thedetector and digital processing unit again similarly co-operate toacquire and store a set of digital information representing the imageproduced by the higher energy pulse. The values obtained representingthe lower energy image are then subtracted from the values representingthe higher energy image.

Since the attenuation of the lower energy x-rays by the soft tissue inthe body is approximately the same as soft tissue attenuation of thehigher energy x-rays, subtraction of the lower energy image data fromthe higher energy image data approximately cancels out the informationdescribing the configuration of the soft tissue. When this informationhas been so cancelled, substantially all that remains in the image isthe representation of bone. In this manner, the contrast and visibilityof the bone is substantially enhanced by energy subtraction.

Energy subtraction has the advantage, relative to temporal subtraction,of being substantially not subject to motion artifacts resulting fromthe patient's movement between exposures. The time separating the lowerand higher energy image acquisitions is quite short, often less than onesixtieth of a second.

Details of energy subtraction techniques in digital radiography andfluoroscopy are set forth in the following technical publications, allof which are hereby incorporated specifically by reference:

Hall, A.L. et al: “Experimental System for Dual Energy ScannedProjection Radiology”. Digital Radiography proc. of the SPIE 314:155-159, 1981;

Summer, F.G. et al: “Abdominal Dual Energy Imaging” Digital Radiographyproc. SPIE 314: 172-174, 1981;

Blank, N. et al: “Dual Energy Radiography: a Preliminary Study”. DigitalRadiography proc. SPIE 314: 181-182, 1981; and

Lehman, L.A. et al: “Generalized Image Combinations in Dual kVp DigitalRadiography”, Medical Physics 8: 659-667, 1981.

Dual energy subtraction has been accomplished, as noted above, bypulsing an x-ray source in a digital scanning slit device at two kVp's,typically 120 and 80 kVp, and sychronizing the pulses with a rotatingfilter which hardens the high kVp pulses by filtering out the lowerenergy x-ray. This results in the patient and x-ray detectorsequentially seeing high energy and low energy beams from which the massper unit area of bone and soft tissue can be solved for.

In energy subtraction, it is desirable that the two energy levels shouldbe widely separated. This is necessary in order to accurately define themasses per unit area of bone and soft tissue.

With a slit scanning device, such as described above, sequentiallypulsing the x-ray tube at 120 and 80 kVp is technically difficult andgives rise to very difficult problems in a practical clinical device.The switching frequency has to be on the order of 500 Hz. andinsufficient photons (x-ray energy per pulse) results when the highestcapacity x-ray tubes are combined with realistically narrow slit widthsand scanning times.

In connection with CT (computerized tomography) applications, a twolayer energy sensitive detector has been proposed. In this proposal, afirst calcium fluoride layer is provided for sensing lower level x-rayradiation, and a second downstream sodium iodide layer senses higherenergy radiation passing through the first layer. Light caused byradiation in each of the two layers is separately sensed by respectivephotomultiplier tubes.

DISCLOSURE OF THE INVENTION

The disadvantages and problems of the prior art are alleviated oreliminated by the use of an energy discriminating radiation detectorincluding three elements. The detector includes a first elementpredominantly responsive to radiation of a first energy range, and asecond element positioned behind the first, responsive to radiation in asecond and higher energy range, along with a radiation filter interposedbetween the first and second elements.

Thus, an energy sensitive x-ray detector system for use in digitalradiography is provided. For each picture element of the radiographicprojection, the detector provides two readings from which the mass perunit area of bone and soft tissue through which the x-ray beam passescan be determined.

The energy sensitive x-ray detector employs a low atomic number phosphorscreen or discrete array of phosphor segments coupled to a photodiodearray, followed by a high atomic number of phosphor screen or discretesegment array similarly coupled.

An energy sensitive segment of an element of the detector systemconsists of a low atomic number phosphor coating layer coupled to afirst photodiode, followed by a high atomic number phosphor coatinglayer coupled to a second photodiode. The low atomic number phosphorpreferentially absorbs the low energy photons emerging from the patientand transmits most of the higher energy photons, a larger percentage ofwhich are absorbed in the second (higher atomic number) phosphor.

Placing an appropriate filter between the two phosphor/photodiode arraysincreases or hardens the effective energy of the x-ray spectrum incidenton the second phosphor and results in a greater and more desirableenergy separation between the x-ray spectra absorbed in the two phosphorlayers.

In accordance with another embodiment, a split energy radiation detectoris provided including a first energy responsive element comprising aquantity of phosphor material including one of yttrium oxysulfide andzinc cadmium sulfide, and a second energy responsive element positionedto receive energy passing through said first element, said secondelement including one of gadolinium oxysulfide and cadmium tungstate.

In accordance with another specific aspect of the invention, theradiation filter interposed between the two elements or layers is madeof a material containing copper.

In accordance with a broader aspect of the invention, there is provideda split energy radiation detector screen comprising a deck of separatedetector elements at least partially mutually superposed, each elementbeing capable of producing information spatially locating radiationincident on the screen.

These and other aspects of the present invention will become moreapparent from a consideration of the following description and of thedrawings, in which:

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 is a plan pictoral and block illustration of a systemincorporating the present invention;

FIGS. 1A-1E are detail views illustrating a portion of the system ofFIG. 1;

FIG. 2 is a side view illustrating a portion of the system illustratedin FIG. 1;

FIG. 2A is a detailed side view illustrating a portion of the system ofFIG. 1;

FIG. 3 is a perspective view of an alternate embodiment of a portion ofthe system of FIG. 1.

FIG. 3A is a graphical description of a preferred feature of the portionof the system illustrated in FIG. 2.

FIG. 4 is a graphical representation of operating characteristics of theportion of the system illustrated in FIG. 2;

FIG. 5 is a block diagram illustrating another system incorporating anembodiment of the present invention.

BEST MODE FOR CARRYING OUT THE INVENTION

FIG. 1 illustrates a slit projection type of digital radiography systemin which the present invention is incorporated. The system S scans athin fan beam of multi-energetic x-rays over a patient's chest andseparately detects a pattern of x-rays of different energies emergentfrom the patient's body. Information represented by the detected x-raysis processed and displayed to illustrate a representation of an image ofthe patient's internal body structure or condition.

More specifically, the system S includes an x-ray source X affixed tomounting structure M for projecting a thin fan beam B of x-rays throughthe body of a patient P, to strike an aligned array D of detectorsegments. The fan beam B is confined by a forward slit K tosubstantially a vertical plane. The detector array D constitutes avertical stack of individual detector segments E, described in moredetail below, and aligned with the vertical plane defined by the beam B.An aft slit J attached to the detector D receives and aids in thedefinition of the planar beam B.

The x-ray source X is mounted on the structure M to rotate about avertical axis, defined in FIG. 1 as extending into the paper. Mechanicallinkage L couples the x-ray tube X to the detector array D and causesthe detector array D to scan behind the patient's body in the directionsof the arrows A, A¹, in order to maintain the detector D aligned withthe beam B throughout the scanning rotative motion of the x-ray tube X.

The x-ray source X is controlled to emit either a continuous beam or arapid succession of x-ray pulses in the form of the fan beam B. Thex-ray tube X and the detector D are synchronously scanned, about avertical axis, across the patient from one side of his body to theother. The detector output is periodically sampled. Each samplingproduces signals representing a line of image information. Over thecourse of the scan from side to side, signals are developed describing aplurality of lines, which together constitute an area image of thepatient's internal body structure.

Details of some aspects of a digital radiography system such asdescribed above are set forth in the following publications, herebyexpressly incorporated by reference.

Arnold, B.A. et al, “Digital Radiography: An Overview” Proc. of S.P.I.E.Vol. 273, Mar. 1981;

Kruger, R.A. et al, “A Digital Video Image Processor for Real Time X-RaySubtraction Imaging” Optical Engineering Vol. 17 No. 6 (1978).

The detector D separately detects x-rays of different energy rangesimpinging on the detector array. An element of the detector array, byway of two sets of leads 01, 02, transmits analog signals representingdetected x-rays within lower and higher energy ranges, respectively.

The signals on the lead sets 01, 02, are provided to ananalog-to-digital converter C which digitizes the outputs and feeds themto a digital processing and receiving unit DPU. The DPU processes thesedigitized output signals to construct a digital representation of animage of the patient's internal body structure scanned by the x-ray beamB, on a line-by-line basis. Digital signals from the DPU are convertedto analog form by way of a digital-to-analog converter, and fed to adisplay unit T, which in response, produces an image in visual formcorresponding to the image representing signals from the DPU.

Optionally, digital storage means can be provided in conjunction withthe DPU in order to digitally store the image representations for futureuse. In such event, the digitally stored signals can be played blackthrough the DPU, converted to analog form, and their correspondingimages displayed at a later time on the display apparatus T.

FIGS. 1A and 1B illustrate (in simplified form, for clarity) particularconfigurations of the face of the detector array D, as viewed from theright in FIG. 1. In FIG. 1A, for example, it is seen that the detector Dcomprises a linear vertically stacked elongated array of detectorsegments E.

An alternative embodiment to the vertical linear detector array shown inFIG. 1A is illustrated in FIG. 1B. This is known as a “staggered” array.The staggered array consists of two side-by-side vertical columns ofdetector segments E, E¹. One of the vertical columns, however, isslightly vertically displaced with respect to the other, by a distanceequal to one-half the height of a single detector segment.

FIGS. 1C-1E illustrate in simplified form several embodiments of thedetector configuration of FIG. 1A as viewed from the right side in FIG.1A. FIGS. 1C-1E, however, are not intended to show the dual layeredstructure of the detector segments, which will be later discussed indetail, such as in connection with FIG. 2A. The detector arrays aredivided into individual segments in one of three ways. In oneembodiment, shown in FIG. 1C, the detector array D comprises anelongated vertical screen strip 10 of particles of radiation sensitivematerial which are glued together with a binder and affixed to a backingof a suitable material such as polyester. The radiation sensitivematerial respnds to incident radiation to produce light. Behind theradiation sensitive screen 10 is a vertical array of adjacentphotodiodes 12. Each photodiode responds to radiation-caused light inthe screen 10 to produce an analog electrical signal indicatingbrightness of the flash caused by the sensed radiation events. Each ofthe photodiodes 12 responds primarily to light from radiation eventsoccurring within a portion of the screen material 10 located adjacentthe photodiode.

Special “cellularized” detector configurations are illustrated in FIGS.1D and 1E. Cellularized detectors have the advantage of reducing theeffects of energy scatter within the detector array.

In the form illustrated in FIG. 1D, the detector screen 10 is grooved asillustrated for example at reference character 14, and the grooves areimpregnated with a reflective material, such as aluminum oxide, toreduce the effects of light within the screen 10. The grooves arealigned with the junctions between each of the adjacent photodiodes 12.

Another form of cellularized detector arrangement is illustrated in FIG.1E. In that embodiment, rather than utilizing an homogeneous screen,with or without grooves, separate crystalline portions 16 of radiationsensitive material are employed. Each crystal is matched to an adjoiningphotodiode and separated from adjacent crystals by a reflective layer.The size of each of the crystals corresponds to the size of itsadjoining photodiode 12.

In all of the foregoing detector arrangements, the photodiodes areadhered to the screen portion 10 by a mechanical pressing operation,which can optionally be aided by a small quantity of adhesive, and/or asmall amount of optical coupling grease to enhance the degree of opticalcoupling between the screen 10, or crystals 16, and the photodiodes 12.

As pointed out above, it is desirable, when practicing the energysubtraction image processing technique, to be able to separatelyrepresent different energy radiation which impinges on the detectorsegments. Herein is disclosed a particular dual layered, energydiscriminating structure for each detector segment which facilitatesachievement of this goal.

FIG. 2 illustrates a particular layered detector segment structure foruse as a component of an energy sensitive radiation detector array D.The detector responds to radiation incident upon it, transmitted in adownward direction with respect to FIG. 2, to produce two outputs atleads 18, 20. The output at lead 18 represents radiation incident uponthe detector segment having an energy level in a lower energy range. Theoutput at the lead 20 represents the detector segment's response toincident x-ray radiation having an energy level in a second, higherenergy range.

The detector segment includes a first elemental layer 22 primarilyresponsive to lower energy x-rays, and a second elemental layer 24responsive to higher energy x-rays. Each of the layers 22, 24, includesa phosphor coating layer 26, 28, respectively, and a photodiode 30, 32,each respectively optically coupled to the phosphor layers 26, 28.

The choice of materials for the phosphor layers 26, 28, is significant.For example, preferred phosphor material for the first phosphor layer 26include yttrium oxysulfide, and zinc cadmium sulfide. Alternativephosphors are barium sulfate, barium cadmium sulfate, lanthimumoxysulfide and barium fluorochloride.

For the second phosphor layer 28, preferred phosphors are gadoliniumoxysulfide and cadmium tungstate. Alternative phosphor materials for thephosphor layer 28 include calcium tungstate and barium lead sulfate.

A preferred phosphor coating weight for the first phosphor layer 26 isabout 20 to 100 milligrams (mg) per square centimeter (cm²).

Preferred phosphor coating weights for the second phosphor layer lie inthe range from approximately 50 to 1000 mg/cm².

The detector segment described above as embodying this invention isuseful not only in linear detector element arrays such as used in scanor slit projection radiography, but also in larger area detector screensused in digital radiography systems incorporated divergent, “area” x-raybeams. In the latter case, a phosphor matrix embodying the detector canconsist of either a single integral x-ray intensifying screen, acellularized intensifying screen, or a cellularized matrix of individualphosphor crystals.

The segments have equal square dimensions in each layer.

The dimensions of the individual cell segments, where a cellularizedstructure is used, are equal to the photodiode matrix array spacing,such that each individual photodiode is congruent with its cell segment.

The cell segment dimensions are greater in the second layer of thedetector than in the first. The relationship between cell segmentdimensions in the first and second layers is expressed by the following:

(D2/D1)=(F2/F1)

where

D2=the second detector photodiode, dimension;

D1=the first detector photodiode dimension;

F2=the distance from the x-ray source focal spot to the second detectorlayer 24, and

F1=the distance from the x-ray focal spot to the first detector layer 22(see FIG. 3A for a graphical illustration of these values).

This relation applies irrespective of whether a slit projection or areascreen is employed.

It is desirable that the phosphor material selected for the firstphosphor layer 26 have a primary absorber atomic number lying in therange of 39 to 57. The corresponding desirable atomic number range forthe phosphor materials' primary absorber selected for the second layer28 is 56 to 83.

The capability of the detector structure of this invention todistinguish between incident x-rays of differing energy ranges can beenhanced by the interposition of a filter layer 36 between the first andsecond layers 22, 24. A preferred filter material is one containingcopper, such as brass. A preferred filter thickness, where brass isused, is approximately 0.5 millimeters (mm). The range of practicalbrass filter thicknesses is from about 0.2 mm to about 1.0 mm.Alternative filters can comprise either single or multiple filterelements made of material ranging in atomic number from approximately 24to 58.

When a detector element constructed in accordance with the presentlyindicated preferred embodiment is used, a desirable energy spectrum forthe x-ray source is from about 80 kVp to 150 kVp, or even higher, iftube technology permits.

The degree of spacing between the first and second layers 22, 24 of thedetector segment is not particularly critical. Spacing between the firstand second layers can suitably vary from almost physical contact toabout 3 or more centimeters (cm). The spacing between the filter layer36 and the first and second layers 22, 24 is not critical either.

As mentioned above, figures such as FIG. 1C show a side view of thedetector array D in a form simplified for clarity. FIG. 1C is simplifiedin that it shows only one of the two detector elements or layers whicheach contain a plurality of detector segments as defined by thedimensions of the photodiodes 12.

FIG. 2A is provided to show the dual detector element (layer) structurewhich is the present subject. FIG. 2A shows how the detailed structureof FIG. 2 appears, when incorporated into a linear detector array D.FIG. 2A represents a side view of such an array.

FIG. 2A illustrates the two detector elements or layers 22, 24 onepositioned behind the other with respect to the incident radiation fromthe source. Each element includes respectively a coating layer ofphosphor 26, 28, and a set of photodiodes respectively indicated at 30,32. Between the elements is located the filter element 36.

Each photodiode has a lead emergent therefrom for transmitting itsanalog radiation indicating signal to the appropriate one of the leadgroups 01, 02, as described generally above. For purposes of clarity,only representative leads are shown in FIG. 2A.

The application of the split energy radiation detector of this inventionis by no means limited to a linear array of detectors, for use in slitprojection digital radiography, the environment described in detailabove. The present invention can also be embodied in a so-called “area”detector, i.e., a relatively large rectangular radiation detectorcovering a relatively expansive portion of the patient's body, designedfor use with so-called “area” beams, which diverge from the source toexpose the radiation detector simultaneously over its entire face. Onelayer of such an area detector is illustrated in FIG. 3, it beingunderstood that such an area detector includes two such layers, onebehind the other.

Other types of area detectors exist in which use of this invention isadvantageous. One such area detector includes a first phosphor layer ofrelatively low atomic number, as described above, coupled to aradiographic film layer, behind which is a second higher atomic numberphosphor screen coupled to a second piece of film. Also, instead of thefilm portions, photoconductive or thermoluminescent plates could beused.

The principles analogous to the construction of the cellularized anduncellularized detectors described above in conjunction with FIGS. 1Athrough 1E can also be applied to area detectors as well.

Where such an area detector is used, the decoding electronics forlocating the sites of radiation events across the face of an areadetector are more complicated than in the case of the linear detectorarray discussed above. Details of a system for accomplishing this, whichcould analogously be applied to an area detector embodying thisinvention, are set forth in publication entitled “A Practical Gamma RayCamera System Using High Purity Germanium” published in the February1974 issue of IEEE Trans Nuc Sci and prepared by the Ohio StateUniversity Department of Nuclear Engineering under the auspices of aNational Institute of Health contract. This publication is expresslyincorporated by reference herein.

As may be implied by the above incorporated publication, the presentinvention is applicable to radiation detector technology employing otherthan phosphor materials which convert radiation events into lightenergy. The principles of this invention can be incorporated as wellinto radiation detection technology utilizing other types of radiationsensitive material, such as solid state materials which convert incidentradiation into electrical signals which represent radiation incident onthe material, without the need for converting such energy to the form oflight.

Energy Sensitive Experiment and Results

The arrangement of the first and second detector layers employed in theexperiment was in effect as shown in FIG. 2. A Lucite and aluminumphantom 38 was employed to simulate soft tissue and bone. Theexperimental results are tabulated in Table 1 for a typical 120 kVpradiation level and plotted in FIG. 4. Note how the iso-Lucite andiso-aluminum lines are more distinct when the brass filter is insertedbetween the first and second detector layers. From the data in Table 1the relative uncertainty in estimating the thickness of Lucite andaluminum can be calculated and these results are tabulated in Table 2.Note that the ability to discriminate Lucite and aluminum is improvedwhen the brass filter is inserted between the first and second detector.

The first phosphor layer was a 43 mg/cm² coating of yttrium oxysulfide.The second phosphor layer was a 110 mg/cm² coating of gadoliniumoxysulfide.

TABLE 1 Experimental Results for a Constant, Typical Exposure LevelBrass Lucite Aluminum (cm) (cm) (cm) (R₁) (R₂) 0 0 0 3167 3809 2.54 01662 2466 5.08 0 917 1451 8.89 0 398 679 10.16 0 309 2.59 529 3.18 11.430 235 415 10.16 .1  275 491 10.16 .2  249 455 10.16 .4  209 2.22 4002.40 10.16 .8  150 308 .0558 0 0 3196 2293 2.54 0 1697 1390 5.08 0 945338 8.89 0 408 400 10.16 0 312 2.47 316 3.06 11.43 0 242 249 10.16 .1 282 298 10.16 .2  255 278 10.16 .4  211 2.21 248 2.29 10.16 .8  154 197

TABLE 2 Lucite and Aluminum Discrimination for 10.2 cm of Lucite and 4mm of aluminum Brass filter Lucite Aluminum % Lucite % AluminumThickness Resolution Resolution Resolution Resolution   0 mm 0.24 cm0.102 cm  2.4 19.0 0.56 cm 0.16 cm 0.05 cm 1.6 12.5

A split energy level radiation detector such as illustrated in detail inFIG. 2 is also applicable in conventional radiography systems as aphototimer. FIG. 5 illustrates such a system. An x-ray source 50 directsa beam 51 of x-ray through the body of a patient P and onto aconventional radiation screen 52. A split level radiation detector 54,constructed in accordance with the structure detailed in FIG. 2 ispositioned as a phototimer behind the screen to receive that portion ofthe x-ray energy from the beam 51 which passes through the screen 52.

The phototimer 54 produces, on leads 53, 55, signals indicating theamount of received energy in separate lower and higher energy ranges,respectively. These separate energy indicating signals are fed to a duallevel energy integrator 56.

The energy integrator 56 includes circuitry for separately integratingthe amount of energy, over time, indicated by the outputs on the leads53, 55.

When the integrated energy values developed by the integrator 56accumulate to a predetermined criteria, the integrator 56 produces asignal to a tube control circuit 58 which terminates operation of thesource 50 in response to the accumulation of the particularpredetermined integrated energy criterian.

The energy criterian governing the time of x-ray exposure can beselected in accordance with known principles by those with skill in theart. This criterion can be defined as the accumulation of apredetermined amount of energy in either of the sensed energy ranges, orcan be a function of both sensed energy levels.

It is to be understood that this description of one embodiment of thepresent invention is intended as illustrative, and not exhaustive, ofthe invention. It is to be further understood that those of ordinaryskill in the relevant art may make certain additions, deletions andmodifications to this embodiment of the invention as described herein,without departing from the spirit or the scope of the invention, asdescribed in the appended claims.

I claim:
 1. An In an imaging system, an energy discriminating radiationdetector comprising: (a) a first element comprising a first material ofa kind which is preferentially responsive to penetrative radiation of afirst energy range; (b) a second element comprising a second materialdifferent in kind from said first material and of a kind which ispreferentially responsive to penetrative radiation of a second energyrange extending higher than said first energy range and which ispositioned to receive radiation which has penetrated through a portionof said first element; and (c) a filter of penetrative radiationinterposed between said first and second elements; and (d) means coupledto said elements for producing an image of a portion of an object fromradiation emerging from the object and incident on the first and secondelements.
 2. The detector of claim 1, wherein said filter containscopper.
 3. The detector of claim 1, wherein said filter comprises brass.4. The detector of claim 2 or 3, wherein said filter is selected to havea thickness of from about 0.2 mm to about 1.0 mm.
 5. The detector ofclaim 1, An energy discriminating radiation detector comprising: (a) afirst element comprising a first material of a kind which ispreferentially responsive to penetrative radiation of a first energyrange; (b) a second element comprising a second material different inkind from said first material and of a kind which is preferentiallyresponsive to penetrative radiation of a second energy range extendinghigher than said first energy range and which is positioned to receiveradiation which has penetrated through a portion of said first element;and (c) a filter of penetrative radiation interposed between said firstand second elements; wherein each said element comprises: (a) a phosphorlayer, and (b) a photodiode optically coupled to the phosphor.
 6. Asplit energy radiation detector comprising: (a) a first energyresponsive element comprising a layer of phosphor material including oneof yttrium oxysulfide and zinc cadmium sulfide; and (b) a second energyresponsive element positioned to receive energy penetrating through saidfirst element, said second element including a second phosphor layercomprising one of gadolinium oxysulfide and cadmium tungstate.
 7. Thedetector of claim 6 further comprising: a copper containing filterelement interposed between said first and second elements.
 8. Thedetector of claim 6, wherein: (a) said first phosphor layer has acoating weight of about 20 to 100 mg/cm², and (b) said second phosphorlayer has a coating weight of about 50 mg/cm² to 1000 mg/cm².
 9. Adigital radiography system comprising: (a) an x-ray source for directingx-rays along a path; (b) a split energy radiation detector spaced fromthe source to receive x-rays from said source, said detector comprising:(i) a first element comprising a first material of a kind which ispreferentially responsive to radiation of a first energy range and beinglocated in said path; (ii) a first sensor for sensing radiation responseof said first element; (iii) a second element at least partiallypositioned to receive source radiation passing through said firstelement, said second element comprising a second material of a kindwhich is preferentially responsive to radiation of a second energy levelextending higher than said first range; (iv) a second sensor for sensingradiation response of said second element; and (c) interpretivecircuitry coupled to said sensors for at least partially digitizinginformation from said sensors and producing from said digitizedinformation a representation of at least a portion of internal bodystructure of a subject when interposed in said path.
 10. The system ofclaim 9 14, wherein said first material includes one of yttriumoxysulfide and zinc cadmium sulfide.
 11. The system of claim 9 14,wherein said second material includes one of gadolinium oxysulfide andcalcium tungstate.
 12. The system of claim 9 14, further comprising: anx-ray filter layer between said first and second elements.
 13. Thesystem of claim 12, wherein said filter layer contains copper.
 14. Thesystem of claim 9, A digital radiography system comprising: (a) an x-raysource for directing x-rays along a path; (b) a split energy radiationdetector spaced from the source to receive x-rays from said source, saiddetector comprising: (i) a first element comprising a first material ofa kind which is preferentially responsive to radiation of a first energyrange and being located in said path; (ii) a first sensor for sensingradiation response of said first element; (iii) a second element atleast partially positioned to receive source radiation passing throughsaid first element, said second element comprising a second material ofa kind which is preferentially responsive to radiation of a secondenergy level extending higher than said first range; (iv) a secondsensor for sensing radiation response of said second element; and (c)interpretive circuitry coupled to said sensors for at least partiallydigitizing information from said sensors and producing from saiddigitized information a representation of at least a portion of internalbody structure of a subject when interposed in said path; wherein saidsensors each comprise a photodiode.
 15. The system of claim 9, A digitalradiography system comprising: (a) an x-ray source for directing x-raysalong a path; (b) a split energy radiation detector spaced from thesource to receive x-rays from said source, said detector comprising: (i)a first element comprising a first material of a kind which ispreferentially responsive to radiation of a first energy range and beinglocated in said path; (ii) a first sensor for sensing radiation responseof said first element; (iii) a second element at least partiallypositioned to receive source radiation passing through said firstelement, said second element comprising a second material of a kindwhich is preferentially responsive to radiation of a second energy rangeextending higher than said first range; (iv) a second sensor for sensingradiation response of said second element; and (c) interpretivecircuitry coupled to said sensors for at least partially digitizinginformation from said sensors and producing from said digitizedinformation a representation of at least a portion of internal bodystructure of a subject when interposed in said path; wherein saidsensors each comprise a photodiode; and wherein said x-ray source iscapable of simultaneously producing x-rays in both said energy ranges.16. The system of claim 9 14, wherein each of said elements issubstantially planar, one said element being substantially behind theother with respect to the source.
 17. An imaging method comprising thesteps of: (a) directing x-rays through a subject to be imaged, saidx-rays including both higher and lower energy radiation; (b) separatelydetecting higher and lower energy x-radiation emergent from the subjectby passing said radiation successively through scintillators comprisingrespectively different kinds of materials each preferentially responsiveto radiation of a different one of said lower and higher energy ranges,including sensing responses of said scintillators; (c) at leastpartially digitizing information derived in said detecting step; (d)processing said digitized information; and (e) utilizing said processeddigital information to produce a representation of internal structure ofthe subject.
 18. The method of claim 17, wherein said digital processingstep includes a step of subtracting information obtained in said lowerenergy sensing step from information obtained in said higher energysensing step.
 19. The method of claim 17, wherein said sensing stepcomprises producing information in response to radiation incident on aplurality of separate detector elements, said information includingspatial location representation of said incident radiation with respectto a said sensing element.
 20. An In an imaging system, an energydiscriminating radiation detecting method utilizing first and seconddetector elements, a first of said elements being preferentiallyresponsive to radiation of a first energy range, a second of saidelements being preferentially responsive to energy radiation of a secondenergy range extending higher than said first energy range, said methodcomprising the steps of; : (a) directing radiation extending over bothsaid first and second energy ranges through a subject; (b) positioningsaid first element to receive incident radiation emergent from thesubject for response thereto; (c) positioning said second element toreceive radiation from the source passing through said first element,and (d) filtering radiation transmitted through said first element priorto the arrival of said energy incident upon said second element; and (e)producing an image of a portion of the subject from the radiationemerging from the subject and incident on the first and second elements.21. A radiographic system comprising: (a) an x-ray source; (b) aradiation detector positioned to receive x-rays from the source; (c) aphototimer comprising: (i) an energy discriminating detector located toreceive x-rays from the source and to produce signals indicating x-rayenergy received in each of two energy ranges, and (ii) circuitry coupledbetween the discriminating detector and the source for controlling thesource as a function of the x-rays detected in said two energy ranges.22. A radiation imaging system comprising: (a) a source of penetrativeradiation; (b) a dual energy detector assembly comprising twoside-by-side columns of individual detector elements, one column beingstaggered with respect to the other by a distance equal to less than thedimension of a single detector element taken along the direction of itscolumn, and additional detector elements positioned behind said columns,relative to said source; (c) mounting structure for maintaining saidsource and said detector assembly sufficiently spaced to provide asubject examining space and for maintaining said detector alignedcontinuously in said penetrative radiation when produced by said source;(d) power means for actuating said source to direct penetrativeradiation through the subject examination space and incident onto thedetector assembly; (e) means coupled to said detector elements forproducing an image of a portion of a subject, when located in thesubject space, from radiation emergent from said subject.
 23. The systemof claim 22, wherein said staggered columns of detector elements areoffset with respect to one another by a distance equal approximatelyone-half the height of a single detector element taken in a directionalong its column.
 24. An energy discriminating radiation detectorcomprising: (a) a first component comprising a first material of a firstkind which is preferentially responsive to penetrative radiation of afirst energy range; (b) a second component comprising a second materialdifferent in kind from said first material and of a kind which ispreferentially responsive to penetrative radiation of a second energyrange extending higher than said first energy range, said secondcomponent being positioned to receive radiation which has penetratedthrough a portion of said first component, and (c) means coupled to saidfirst and second components to produce electrical signals representingradiation when incident respectively on said first and secondcomponents.
 25. The detector of claim 24, wherein: said filter comprisesfurther comprising a filtering material having an atomic number in therange of 24-58.
 26. The detector of claim 24, wherein: said firstcomponent comprises a phosphor layer comprising an element having anatomic number lying in the range of 39-57.
 27. The detector of claim 24,wherein said second component comprises: a phosphor layer comprising anelement having an atomic number lying within the range of 56-83.
 28. Thedetector of claim 24, wherein one of said first and second componentscomprises: a phosphor layer proximate and aligned with a layer of lightsensitive film.
 29. The detector of claim 24 1, wherein one of saidfirst and second componentselements comprises: a phosphor layer; andwherein said means for producing an image comprises a photoconductiveplate, said phosphor layer being proximate and aligned with a portion ofsaid photoconductive plate.
 30. The detector of claim 24 1, wherein oneof said first and second componentselements comprises: a phosphor layer;and wherein said means for producing an image comprises athermoluminescent plate, said phosphor layer being proximate and alignedwith a portion of said thermoluminescent plate.
 31. The detector ofclaim 24, further comprising: a filter of said penetrative radiationinterposed between said first and second components.
 32. The detector ofclaim 31, wherein: said filter comprises material having an atomicnumber in the range of 24-58, and a thickness in the range of about 0.2mm to 1.0 mm.
 33. the detector of claim 24, wherein said second materialcomprises material having a primary radiation absorber having a higheratomic number than that of said first material.
 34. The detector ofclaim 24, wherein: (a) said first material comprises one of yttriumoxysulfite oxysulfide, zinc cadmium sulfide, barium sulfate, bariumcadmium sulfate, lanthium oxysulide lanthanum oxysulfide and bariumfluorochloride, (b) said second material comprises one of gadoliniumoxysulfide, cadmium tungstate, calcium tungstate and barium leadsulfate.
 35. The detector of claim 24, wherein: (a) said first materialcomprises a first layer of phosphor material having a coating weight ofabout 20 to 100 mg/cm², and (b) said second material comprises a secondphosphor layer having a coating weight of about 50 mg/cm² to 1000mg/cm².
 36. The detector of claim 24, wherein: (a) said first componentcomprises a portion of a first scintillator material, and (b) saidsecond component comprises a portion of a second scintillator material.37. The detector of claim 24, further comprising: a portion ofpenetrative radiation filtering material interposed between said firstand second components and being capable of absorbing substantially allradiation incident on said filter element lying within said first energyrange, while not absorbing substantially all such radiation of saidsecond energy range.
 38. A method for detecting area distribution ofdiffering energy levels of penetrative radiation, said methodcomprising: (a) detecting preferentially lower energy radiation bypassing it through a first detector element including a scintillator anda plurality of segments; (b) detecting higher energy radiation bytransmitting radiation emergent from said first detector incident onto asecond detector element including a scintillator and a plurality ofsegments; (c) filtering penetrative radiation emergent from said firstdetector before said second detecting step, and (d) producinginformation in said first and second detecting steps spatially locatingradiation over an area with respect to at least one of said detectorelements.
 39. An energy discriminating radiation detector comprising:(a) a first component comprising a first phosphor material including aprimary radiation absorber having an atomic number lying the range of39-57; (b) a second component comprising a second phosphor materialaligned with said first phosphor material to receive radiation when saidradiation has penetrated through a portion of said first component, saidsecond phosphor material including a primary radiation absorber havingan atomic number lying within the range of 56-83, and (c) means coupledto said first and second components for producing electrical signalsrepresenting radiation when incident on said detector.
 40. An energydiscriminating radiation detector comprising: (a) a first componentcomprising a first material of a first kind which is preferentiallyresponsive to penetrative radiation of a first energy range; (b) asecond component comprising a second material different in kind fromsaid first material and of a kind which is responsive to penetrativeradiation of a second energy range extending higher than said firstenergy range, said second component being aligned with said firstcomponent to receive radiation when said radiation has penetratedthrough a portion of said first component, and (c) means coupled to saidfirst and second components to produce electrical signals representingpenetrative radiation when incident on said detector.
 41. A radiationimaging system comprising: (a) a source for propagating penetrativeradiation along a path from a focal spot; (b) a detector assembly spacedfrom said source and interposed in said path, said detector assemblycomprising: (i) a front array of individual detector elements, eachfront array element including a penetrative radiation sensitivereceiving face having a discrete geometry, said front array elementfaces being located at substantially a distance F₁ from said focal spot;(ii) a rear array of individual detector elements, each said rear arrayelement including a penetrative radiation sensitive receiving facehaving a discrete geometry, wherein each element of said rear array issubstantially aligned behind a corresponding element of said frontarray, with respect to said focal spot, and in which each rear arrayelement has a receiving face which has a larger area then the receivingface of its corresponding aligned front array element, said rear arrayreceiving faces being located at substantially a distance F₂ from saidfocal spot, and (c) circuitry coupled to said detector arrays forproducing a representation of radiation when incident on said detectorelements.
 42. The system of claim 41, wherein: (a) said receiving facesof said detector elements of said front and rear arrays have similargeometry, and (b) a dimension D₁ of one of said front array elements isrelated to a corresponding dimension D₂ of one of said rear arrayelements by the following relation: D₂/D₁=F₂/F₁.